The basic NMR phenomenon is the property of all nuclei having an odd number of protons and/or neutrons to act as small magnets. In the absence of an applied magnetic field, the magnetic axes of these nuclei point in random directions; however, when placed in a strong applied static magnetic field, these axes tend to align with the field. If radio frequency energy of the proper resonant frequency and having its r.f. magnetic field component perpendicular to the static magnetic field is then transmitted throughout the object under test, the resonant nuclei will nutate (e.g. turn) their magnetic axes (e.g., "flip") against the applied static magnetic field. If such radio waves are then terminated (e.g., such as by termination of an appropriate RF pulse), the magnetic axes of the earlier "flipped" nuclei tend to relax back toward their original alignment with the applied static magnetic field. In doing so, radio frequency waves (of the same frequency as that applied in the earlier flipping r.f. pulse if the magnetic field at the nuclei location is unchanged) are emitted and may be detected by an appropriate radio frequency receiver arrangement.
The first and longest NMR r.f. response to appear is the FID (free induction decay). However due to relative dephasing phenomena of the spinning nuclei during the FID it decays and if the dephasing is reversed a so-called spin echo signal can be generated. For example, if a 90.degree. nutation r.f. pulse is first applied, the flipped nuclei will begin to "dephase" in their relative rotational speeds. If after t seconds a 180.degree. nutation r.f. pulse is applied, the dephasing sense of each flipped nuclei is reversed resulting in an "in phase" condition after another t seconds. This "in phase" condition results in a detectable pulse of r.f. energy known as a spin-echo.
Because there are known relationships between the strength of applied magnetic fields and the frequencies of resultant NMR responses, this NMR excitation/detection sequence can be utilized to obtain basic information concerning the location and distribution of specific nuclei within an object under test.
For example, in any given magnetic field, the frequency of transmitted RF energy required to product NMR is specific (hydrogen in a magnetic field of about 3.5 KG will resonate at about 15 MHz). For all types of nuclei, the so-called Larmor resonant frequency changes in direct ratio to changes in the strength of the surrounding field at the nucleus site. For instance, hydrogen in a magnetic field of about 7 KG exhibits NMR at about 30 MHz. The constant of proportionality between NMR frequency and the instantaneous magnetic field strength at the location of the nucleus is called the magnetogyric ratio and each specific nucleus having an odd number of protons and/or neutrons has its own respectively corresponding magnetogyric ratio constant.
Actually, it is believed that only a very small fraction (e.g., two or three parts per million) of the relevant nuclei within a given measurement volume actually generate the observed NMR phenomenon at any given instant of time. This is apparently a statistical process and the actual nuclei being observed will change over time but sufficient numbers of nuclei are observed at any given instant of time so as to permit significant NMR measurements.
The elapsed time required to obtain maximum alignment of the nuclei magnetic axes with an imposed static magnetic field is typically on the order of one second for hydrogen in tissue. This exponential NMR alignment time is normally denoted by its exponential time constant "T1" (i.e. the time required to obtain 1-l/e of the asymtote or final expected value) and is sometimes known as the longitudinal magnetic relaxation time or as the spin-lattice magnetic relaxation time. It is a function of many local physical and chemical factors including molecular structure, elemental composition, temperature and viscosity. In general, even if only hydrogen nuclei are observed in an NMR imaging scanner, and even if it is assumed that all tissues have equal hydrogen densities, the measured T1 NMR parameter may be expected to differ significantly between different body tissues.
The rate at which NMR signal emission decays is another characteristic exponential time factor and is usually less than the T1 value. This second NMR time factor is commonly referred to by its exponential time constant "T2" and is sometimes known as the transverse magnetic relaxation time or spin-spin magnetic relaxation time. It constitutes another NMR time parameter that is dependent not only upon local physical and chemical factors including molecular structure, elemental composition, temperature and viscosity (not necessarily in exactly the same way as is the T1 parameter however). Accordingly, the T2 parameters are in general also different for different body tissues. For example, the nuclei of very pure liquids, in general, align with an applied static magnetic field less quickly and emit NMR signals for a longer time than do nuclei of liquids loaded with proteins.
Since the time constants associated with NMR phenomena are quite long compared to readily achievable response times of electronic circuits (i.e., radio frequency and magnetic gradient coils), it is possible to use a succession of different magnetic gradients and RF pulses to selectively produce NMR signal responses that can be detected and associated with specific elemental internal volumes of an object under test.
In general, since the frequency of the RF energy required to excite NMR and/or of resultant NMR signals is proportional to the instantaneous magnetic field strength at the measured volume, if the magnetic field strength has a spatial distribution that is known, then the frequency spectrum of NMR exciation/detected signals also encodes the spatial distribution of the NMR nuclei.
Varying the elapsed time interval between successive excitations of a measured volume will produce different amplitudes of NMR response signals in accordance with the T1 parameters associated with the nuclei of the measured volume. That is, if the interval between successive NMR excitations is relatively short, tissues with longer T1 parameter values will yield relatively less NMR response signal than those with shorter T1 parameter values since the former have less chance to become fully re-aligned with the static magnetic field before a new measurement cycle is initiated. In addition, varying the elapsed time interval between the initiation of a measurement cycle (i.e., the first NMR excitation RF pulse required to eventually result in a desired NMR response signal) and the subsequent occurrence of the desired NMR spin echo response signal will produce different corresponding amplitudes of NMR spin echo response in accordance with the T.sub.2 parameters of the tissues in the measured volume. That is, tissues in the measured volume having longer T2 parameter values will provide relatively larger NMR response signals than those having shorter T2 parameter values.
Motion factors can also change the resultant NMR signal intensity in a live object. For example, if hydrogen nuclei move through the measurement volume during one measurement cycle (e.g., approximately 35 milliseconds), the potential NMR response signal from these nuclei will be completely lost. On the other hand, if only a fraction of such nuclei remain within the measurement volume during the measurement cycle time, the intensity of the NMR response signal will be correspondingly reduced. As should be apparent, the actual reduction in intensity of NMR response signals due to motion factors depends upon the fraction of effected nuclei that are in motion and upon their velocity.
Accordingly, images of an object cross-section constructed from measured intensity of NMR responses represent a complex function of physical characteristics of the tissue and of selectable instrument parameters which can be selectively manipulated.
The instrumentation for NMR imaging systems reflects the sequence by which nuclear magnetic resonance is achieved. A typical system will include a large magnet to create the surrounding magnetic field, magnetic field gradient producing coils to create position dependent magnetic fields, an RF coil to apply and receive the resonant frequency r.f. signals, electronic circuitry to generate, transmit and record the electromagnetic radiations, and a digital data acquisition, processing, and display system.
A variety of different NMR methods to define a measured volume have been developed. All techniques, however, are based on the relationship between RF frequency and magnetic field. Because it is impossible to create a magnetic field with a different strength at every point in space at the same time, all techniques use changing magnetic fields to define volumes. Magnetic field gradients can be used during transmission or reception or both.
The exemplary NMR imager described in the earlier-referenced applications and patents illustrates electronic selection of desired sub-volumes in the object. A slice in the sample can be exited by exposing the sample to a magnetic field variation and a r.f. magnetic field such that only the desired plane corresponds to the frequencies of the r.f. magnetic field. Two intersecting such planes (actually planar volumes) can be excited and the two r.f. fields which separately excite the planes can be chosen so that a signal known as a spin echo will be emitted (at a later time) only by nuclei located at the intersection of the two earlier excited planes (which intersection can itself be a planar volume where the two selected planar volumes are coplanar). The spin echo under these circumstances contains information about only the nuclei located within this common intersecting volume. By applying a field variation within such an intersecting volume, for example, during read out, a spatially-encoded frequency spectrum of the emitted r.f. signal is produced. The intensity I of each frequency component of the spin echo will be a function of the hydrogen density H, the T1, and the T2 parameters of the selected volume element. A map of the hydrogen density modified by T1 and T2 can thus be obtained from the frequency spectra of the spin echo signals. The relaxation times cna be measured by observing signal strength when the relevant T 1 and T2 instrument parameters are varied.
In NMR, as already explained, each tissue is characterized by three parameters: Hydrogen density (H), and the rate at which the polarization of the hydrogen nuclei changes, given by the two times, T1 and T2. The imaging procedures used result in data that is dependent in a complex manner on all three parameters.
The relationship of observed NMR intensity (I) to the four physical parameters (H, T1, T2, and motion), is approximately given by: EQU I=Hf(v) exp (-TE/T2)[1-exp (-TR/T1)] Eq. 1
(Note: Actually, Equation 1 includes a denominator of 1+exp (-TR/T2) exp (-TR/T1) which can be assumed as substantially equal to unity for TR greater than about 3T2) where I is the NMR intensity in a particular region of the image; H is the local hydrogen density; TE is the T2 parameter of the instrument (e.g. the time delay of a detected spin echo after a 90.degree. flip pulse and affects T2 contrast), measured, for example, in milliseconds and varied within a broad range (typically 20-60 ms) under computer control; TR is the T1 parameter of the instrument (e.g. the repetition time of a complete measurement sequence and affects T1 contrast), measured, for example, in seconds and also computer controlled (typically 0.25-2.0 seconds); f(v) is a function of both the speed with which hydrogen nuclei move through the region being imaged and of the fraction of the total number of nuclei that are moving.
It is clear that if I could be measured for TR=infinity and TE=0, the result would be Hf(v). In fact, neither of these values can be reached directly. Thus, the only image that can be directly observed is a distribution of I in space. However, obtaining three images (two with different values of the T1 instrument parameter and the T2 instrument parameter held constant, and a third one where, for one value of the T1 instrument parameter, the T2 instrument parameter is changed), arithmetic manipulation can yield images of Hf(v), T1, and T2. However this has in the past required iterative curve fitting techniques.
Keeping the four parameters of the tissue in mind (H, T1, T2, and motion), resulting NMR intensities from different tissues can be understood. Hydrogen density and movement strongly affect the intensity of signal emitted from:
(a) air, becuase it has essentially no hydrogen and bone because what hydrogen it has is in a solid matrix, will show no signal at all (except where partial volume effects come into play); PA1 (b) vessels through which blood flows will yield a low density signal, and organs with high blood flow will show a lower and "blochier" NMR intensity in a live animal when compared to signals that would obtain in the dead animal; and PA1 (c) the lung, because of its air content of approximately 70% and high blood flow, will also be of lower intensity than other soft tissues.
T1 and T2 NMR parameters play an important role in distinguishing between soft tissues. The hydrogen content of most soft tissues varies over a range of approximately 20%. Contrasting with this, T1 and T2 vary by factors of up to 500%. Tissues with short T1 and T2 values will tend to yield the largest NMR intensities, and because of the exponential nature of the modulation of H, the relatively smaller variations of this latter parameter may be overshadowed by T1 and T2 variations.